Stent made from an ultra high molecular weight bioabsorbable polymer with high fatigue and fracture resistance

ABSTRACT

A stent made from an ultra high molecular weight bioabsorbable polymer is disclosed herein. The bioabsorbable polymer can have a Mw greater than 1 million g/mole or greater than 2 million g/mole. Methods of making the ultra high molecular weight polymer stent without degrading the molecular weight are further disclosed.

This application is a divisional application of U.S. patent applicationSer. No. 12/422,783 filed Apr. 13, 2009, which is incorporated byreference herein.

BACKGROUND OF THE INVENTION

Field of the Invention

This invention relates to methods of manufacturing polymeric medicaldevices, in particular, stents.

Description of the State of the Art

This invention relates to radially expandable endoprostheses, that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts.

Stents are often used in the treatment of atherosclerotic stenosis inblood vessels. “Stenosis” refers to a narrowing or constriction of abodily passage or orifice. In such treatments, stents reinforce bodyvessels and prevent restenosis following angioplasty in the vascularsystem. “Restenosis” refers to the reoccurrence of stenosis in a bloodvessel or heart valve after it has been treated (as by balloonangioplasty, stenting, or valvuloplasty) with apparent success. Stentare also used widely in endovascular applications, such as in thepopliteal artery.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, formed fromwires, tubes, or sheets of material rolled into a cylindrical shape.This scaffolding gets its name because it physically holds open and, ifdesired, expands the wall of the passageway. Typically, stents arecapable of being compressed or crimped onto a catheter so that they canbe delivered to and deployed at a treatment site.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance.Effective concentrations at the treated site require systemic drugadministration which often produces adverse or even toxic side effects.Local delivery is a preferred treatment method because it administerssmaller total medication levels than systemic methods, but concentratesthe drug at a specific site. Local delivery thus produces fewer sideeffects and achieves better results.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesan active or bioactive agent or drug. Polymeric scaffolding may alsoserve as a carrier of an active agent or drug.

The stent must be able to satisfy a number of mechanical requirements.The stent must be capable of withstanding the structural loads, namelyradial compressive forces, imposed on the stent as it supports the wallsof a vessel. Therefore, a stent must possess adequate radial strength.Radial strength describes the external pressure that a stent is able towithstand without incurring clinically significant damage. Additionally,a stent should be sufficiently rigid to adequately maintain its size andshape throughout its service life despite the various forces that maycome to bear on it, including the cyclic loading induced by the beatingheart. For example, a radially directed force may tend to cause a stentto recoil inward. Furthermore, the stent should possess sufficienttoughness or resistance to fracture from stress arising from crimping,expansion, and cyclic loading.

Some treatments with implantable medical devices require the presence ofthe device only for a limited period of time. Once treatment iscomplete, which may include structural tissue support and/or drugdelivery, it may be desirable for the stent to be removed or disappearfrom the treatment location. One way of having a device disappear may beby fabricating the device in whole or in part from materials that erodeor disintegrate through exposure to conditions within the body. Thus,erodible portions of the device can disappear or substantially disappearfrom the implant region after the treatment regimen is completed. Afterthe process of disintegration has been completed, no portion of thedevice, or an erodible portion of the device will remain. In someembodiments, very negligible traces or residue may be left behind.Stents fabricated from biodegradable, bioabsorbable, and/or bioerodablematerials such as bioabsorbable polymers can be designed to completelyerode only after the clinical need for them has ended. However, thereare potential shortcomings in the use of polymers as a material forstents such as low fracture toughness.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a stent comprisinga stent body for supporting a vascular lumen, wherein the stent body isformed of a bioabsorbable polymer, wherein the bioabsorbable polymer hasa weight average molecular weight greater than 1 million g/mole.

Additional embodiments of the present invention include a method ofmaking a stent body for supporting a vascular lumen, comprising:immersing a cylindrical member in a solution including a bioabsorbablepolymer dissolved in a fluid, wherein the bioabsorbable polymer has aninherent viscosity greater than 6 dl/g in chloroform at 25° C. or has aweight average molecular weight greater than 1 million g/mole, removingthe member from the solution, wherein a portion of the solution remainson the surface of the member upon removal from the solution; removingsolvent from the solution remaining on the member to form a tubularlayer of the bioabsorbable polymer on the member; repeating theimmersion step, removal from the solution step, and removal of thesolvent step to form a final tubular layer of bioabsorbable polymer onthe member of a desired thickness; and forming a stent body from thefinal tubular layer.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIGS. 2A-B depict a cross-section of a strut of a stent body transverseto its longitudinal axis and along its longitudinal axis, respectively.

FIGS. 3A-B depict a cross-section of a strut of a stent body orscaffolding transverse to its longitudinal axis and along itslongitudinal axis, respectively.

FIG. 4 depicts a transverse cross-section of a strut with a coatinglayer.

FIGS. 5A-C illustrate the dip coating process of the present invention.

FIGS. 6A-B depict radial and axial cross-sections, respectively, of acoated mandrel.

FIG. 7 depicts a mandrel mounting disk having a plurality of holesconfigured to hold mandrels for a dip coating operation.

FIG. 8A depicts a system for controlled dip coating of mandrels mountedon the mounting disk of FIG. 7.

FIG. 8B shows the system of FIG. 8A with the mounting disk and mountedmandrels removed from a solution.

FIG. 9 depicts the results of a dip coating operation.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of the present invention relate to implantablemedical devices, such as stents made from bioabsorbable polymers thatprovide high fatigue and fracture resistance. These bioabsorbablepolymers have an ultra high molecular weight, which provide the highfatigue and fracture resistance as compared to typical bioabsorbablepolymers with much lower molecular weights fabricated by typical meltprocessing techniques such as extrusion or injection molding. Theembodiments further include methods of fabricating the stent. Theembodiments are generally applicable to any tubular polymericimplantable medical device. In particular, the methods can be applied totubular implantable medical devices such as self-expandable stents,balloon-expandable stents, and stent-grafts.

A stent may include a pattern or network of interconnecting structuralelements or struts. FIG. 1 depicts a view of a stent 100. In someembodiments, a stent may include a body, backbone, or scaffolding havinga pattern or network of interconnecting struts or structural elements105. Stent 100 may be formed from a tube (not shown). The structuralpattern of the device can be of virtually any design. The embodimentsdisclosed herein are not limited to stents or to the stent patternillustrated in FIG. 1. The embodiments are easily applicable to otherpatterns and other devices. The variations in the structure of patternsare virtually unlimited. A stent such as stent 100 may be fabricatedfrom a tube by forming a pattern with a technique such as laser cuttingor chemical etching.

A stent such as stent 100 may be fabricated from a polymeric tube or asheet by rolling and bonding the sheet to form the tube. A tube or sheetfor making a stent are conventionally formed by extrusion or injectionmolding. A stent pattern, such as the one pictured in FIG. 1, can beformed in a tube or sheet with a technique such as laser cutting orchemical etching. The stent can then be crimped on to a balloon orcatheter for delivery into a bodily lumen.

An implantable medical device can be made partially or completely from abiodegradable, bioabsorbable, or biostable polymer. A polymer for use infabricating an implantable medical device can be biostable,bioabsorbable, biodegradable or bioerodable. Biostable refers topolymers that are not biodegradable. The terms biodegradable,bioabsorbable, and bioerodable are used interchangeably and refer topolymers that are capable of being completely degraded and/or erodedwhen exposed to bodily fluids such as blood and can be graduallyresorbed, absorbed, and/or eliminated by the body. The processes ofbreaking down and absorption of the polymer can be caused by, forexample, hydrolysis and metabolic processes.

A stent made from a biodegradable polymer is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished. Afterthe process of degradation, erosion, absorption, and/or resorption hasbeen completed, no portion of the biodegradable stent, or abiodegradable portion of the stent will remain. In some embodiments,very negligible traces or residue may be left behind.

The duration of a treatment period depends on the bodily disorder thatis being treated. In treatments of coronary heart disease involving useof stents in diseased vessels, the duration can be in a range from abouta month to a few years. However, the duration is typically up to aboutsix months, twelve months, eighteen months, or two years. In somesituations, the treatment period can extend beyond two years. The stentis expected to be completely degraded away from the vessel at the end ofthe treatment period.

As indicated above, a stent has certain mechanical requirements such ashigh radial strength, high modulus, high fracture toughness, and highfatigue resistance including bending fatigue for endovascularapplications. A stent that meets such requirements greatly facilitatesthe delivery, deployment, and treatment of a diseased vessel. Apolymeric stent with inadequate mechanical properties can result inmechanical failure, strut fracture, or recoil inward after implantationinto a vessel.

With respect to radial strength, the strength to weight ratio ofpolymers is usually smaller than that of metals. To compensate for this,a polymeric stent can require significantly thicker struts than ametallic stent, which results in an undesirably large profile.

Additionally, polymers that are sufficiently rigid to support a lumen atconditions within the human body may also have low fracture toughnesssince they may exhibit a brittle fracture mechanism. For example, theseinclude polymers that have a glass transition temperature (Tg) abovehuman body temperature (Tbody), which is approximately 37° C. Suchpolymers may exhibit little or no plastic deformation prior to failure.It is important for a stent to be resistant to fracture throughout therange of use of a stent, i.e., crimping, delivery, deployment, andduring a desired treatment period. In particular, stents used inendovascular applications, such as in the popliteal, are highlysusceptible to fracture as a result of severe bending and torsionalfatigue that occurs in the peripheral anatomy as a result of walking,running, sitting and other physical activity. PLLA is but one example ofthe class of semicrystalline polymers for which the above description istrue. The Tg of PLLA varies between approximately 50 to 80° C., or morenarrowly between 55 and 65° C., depending on crystallinity,microstructure, and molecular weight.

A tubular polymer perform or construct, such as a tube, sheet, orfibers, from can be made from a polymer melt of semicrystalline polymersfrom methods such as extrusion or injection molding. The constructsproduced from these methods and the devices, such as a stent, madetherefrom are fairly brittle and have a low elongation to fracture. Suchstents can be prone to fracture when they are crimped and deployed.Additionally, such stents may have insufficient strength to adequatelysupport a lumen.

The weight average molecular weight (Mw) of polymers that is typicallyused for forming constructs for fabricating medical devices is in therange of about 100,000 to 600,000 g/mole or have an inherent viscosityof less than about 4 dl/g. In this range, the viscosity of melts of suchpolymers is low enough to allow melt processing at temperatures that donot significantly degrade the molecular weight of the polymer (between20 to 30%). Without further processing or material modification, of suchconstructs, the mechanical properties are limited by the low molecularweight. Additionally, further processing steps, in particular radiationsterilization, can further degrade the molecular weight. For example,for a starting (e.g., an extruded tube) molecular weight, Mw, of 600,000g/mole, the molecular weight can decrease to Mw of 190,000 g/mole aftere-beam sterilization with a dose of 20-30 kGy.

Various embodiments of the present invention include a stent having astent body or scaffolding formed, fabricated from, or consistingessentially of a bioabsorbable polymer having an ultra high molecularweight (UHMW). The UHMW corresponds to a Mw that is greater than 1million g/mole and an Mn greater than 500,000 g/mole. The polymerconstructs and stent made therefrom may be formed from polymers with aMw between 1-1.2 million g/mole, 1.2-1.5 million g/mole, 1.5 to 1.7million g/mole, 1.7-2 million g/mole, 2 million g/mole, or greater than2 million g/mole. Alternatively or additionally, the bioabsorbablepolymer has an inherent viscosity of greater than 8 dl/g in chloroformat 25° C. Polymer constructs with such a molecular weight range may beobtained from solvent processing methods adapted to solutions of ultrahigh molecular weight polymers. The UHMW polymer constructs and stentsformed therefrom have high strength, high stiffness (modulus), highfracture toughness, and good fatigue resistance.

A polymer construct composed of a UHMW as indicated above cannot befabricated using melt processing methods such as extrusion or injectionmolding. Therefore, a stent body composed of a UHMW polymer made bymachining a stent pattern from a tube cannot be fabricated through meltprocessing. The inventors have found that at the temperatures of meltprocessing of polymers with a conventional range of Mw or inherentviscosity, the viscosity of the UHMW polymer melts is too high toprocess. Alternatively, the inventors have also found that meltprocessing the UHMW polymer at higher temperatures to reduce theviscosity of the melt results in significant molecular weightdegradation of the polymer, so that the polymer in the construct that isformed does not have a UHMW.

In some embodiments, the UHMW polymer of the construct and stents madetherefrom is a semicrystalline polymer. In such embodiments, the polymermay have a Tg above Tbody. In order to provide greater stability of thestent at conditions within a human body, the Tg may be at least 10, 20,or 30° C. greater than human body temperature, which is about 37° C. TheTg of this polymer as measured by DSC is approximately 64° C. at a scanrate of 20° C. per minute as measured by taking the half-height in theΔCp region of (change in heat capacity) of the DCS thermogram.

An exemplary UHMW polymer can be UHMW PLLA. The degree of crystallinityof a UHMW PLLA construct can be 10-40%, or more narrowly, 15-25%,however, the degree of crystallinity can be lower than 15% or greaterthan 25% and depends on processing conditions. The degree ofcrystallinity is as measured by DSC and assuming ΔH_(f)° is 93.7joules/gram 100% crystallinity. The Tg of the UHMW PLLA may be betweenabout 63-65° C. and the Tm of the UHMW PLLA may be between 174-178° C.as measured by DSC at a scan rate of 20° C. per minute. If the PLLA isblended with PDLA to form a stereocomplex, then the Tm can be as high as230. A significant advantage of the UHMW polymer is that it is expectedto provide very high fracture toughness and elongation withoutsacrificing modulus and radial strength for a stent with a relativelylow degree of crystallinity (e.g., less than 25%, or more narrowly lessthan 15%, with a lower degree of crystallinity preferred), whichminimizes brittle behavior. The inventor has found that polymers in theconventional or typical molecular weight range should have a degree ofcrystallinity of greater than about 40-50% to provide sufficient radialstrength and modulus for a stent to support a blood vessel (e.g.,support for at least 1-3 months with less than 10% recoil). Such highdegree of crystallinity provides a greater risk of brittle behaviorduring use of a stent.

The body, scaffolding, or substrate of a stent may be primarilyresponsible for providing mechanical support to walls of a bodily lumenonce the stent is deployed therein. A stent body, scaffolding, orsubstrate, for example, as pictured in FIG. 1, can refer to a stentstructure with an outer surface to which no coating or layer of materialdifferent from that of which the structure is manufactured. If the bodyis manufactured by a coating process, the stent body can refer to astate prior to application of additional coating layers of differentmaterial. “Outer surface” refers to any surface however spatiallyoriented that is in contact with bodily tissue or fluids. A stent body,scaffolding, or substrate can refer to a stent structure formed by lasercutting a pattern into a tube or a sheet that has been rolled into acylindrical shape.

In some embodiments, the stent body, scaffolding, struts, or structuralelements of the present invention may be nonporous or substantiallynonporous. Substantially nonporous refers to a porosity of less than 0.1percent. Alternatively, the stent body, scaffolding, struts, orstructural elements of the present invention may be porous.Additionally, the surface of the stent body, scaffolding, struts, orstructural elements of the present invention may have cavities oralternatively, be cavity-free.

Exemplary semicrystalline polymers that may be used in embodiments ofthe present invention include poly(L-lactide) (PLLA), polyglycolide(PGA), polymandelide (PM), polycaprolactone (PCL), poly(trimethylenecarbonate) (PTMC), polydioxanone (PDO), poly(4-hydroxy butyrate) (PHB),and poly(butylene succinate) (PBS). An exemplary amorphous polymer thatmay be used includes poly(DL-lactide) (PDLLA). Additionally, block,random, and alternating copolymers of the above polymers may also beused in embodiments of the present invention, for example,poly(L-lactide-co-glycolide). PLLA, PM, PGA, PDLLA, and PLGA have Tg'sabove human body temperature.

FIGS. 2A-B depict cross-sections of a strut 120 of a stent bodytransverse to its longitudinal axis and along its longitudinal axis,respectively. Strut 120 has an abluminal surface 122, a luminal surface124, and sidewall surfaces 126. Strut 120 has a thickness Ts, whichcorresponds to the thickness of a tubular construct from which the stentbody is machined. An exemplary range of Ts for a polymer stentscaffolding to adequately support a vessel is between 0.004-0.008 in.

In further embodiments, the struts or structural elements of a stent canhave two or more layers transverse to the radial direction of the stent.Radial direction corresponds direction defined by the shortest line froma point on the structural element to the cylindrical axis of the stent.Each of the layers on the sidewalls of the structural element may befree of a material different from a respective layer. In suchembodiments, the adjacent layers are composed of materials withdifferent properties or characteristics. Different properties orcharacteristics can refer to different chemical composition of apolymers, differing composition due to blending of polymers, differingcomposition due to blending of filler materials, molecular weight, or acombination thereof.

The two or more layers can include an abluminal layer, a luminal layer,and one or more layers between the abluminal and luminal layers. FIGS.3A-B depict cross-sections of a strut 130 of a stent body or scaffoldingtransverse to its longitudinal axis and along its longitudinal axis,respectively. Strut 130 has a luminal layer 132 and an abluminal layer134. A surface 139 is a boundary between the two layers. Luminal layer132 has a luminal surface 136 and abluminal layer 134 has an abluminalsurface 138. Abluminal layer 134 has a thickness T1 and luminal layer132 has thickness T2.

In some embodiments, one of the layers is an UHMW polymer and otherlayers can be made of a low molecular weight polymer with a Mw that isin a conventional range, for example, less than 300,000 g/mole. In anexemplary embodiment, the luminal layer can be made of an UHMW polymerand the abluminal layer can be a low molecular weight layer, althoughthe layers could be reversed. The ratio of the thicknesses of the UHMWlayer and the low molecular weight layer can vary, for example, between1000:1 to 2:1. In other embodiments, adjacent layers can be made fromUHMW polymers, but the polymers in the different layers have differentMw.

The use of low molecular weight layer(s) with a UHMW layer allowsadjustment of the degradation time (i.e., time for stent to completelyerode from an implant site) to a desired range. A UHMW polymer layer hasa longer degradation time than a lower molecular weight polymer layer. Astrut will lose structural integrity and disintegrate faster as thethickness of the UHMW is decreased. Additionally, the low molecularweight layer can be a therapeutic layer. The low molecular weightpolymer can have a therapeutic substance dispersed therein. For example,the ratio an UHMW layer to a low molecular weight layer can be 500:1 to1000:1 with the low molecular weight layer being a therapeutic layer.

In further embodiments, the struts or structural elements of thescaffolding or body of the stent made from a UHMW polymer can have apolymer coating layer disposed partly or completely around the surfaceof the strut or structural element. FIG. 4 depicts a transversecross-section of a strut 140 with a coating layer 142 disposed above anabluminal surface 144, luminal surface 146, and sidewall surfaces 148 ofstrut 140. Alternatively, the coating can be disposed exclusively abovethe abluminal layer, luminal layer, or both the luminal and abluminallayers. In such embodiments, the coating includes a polymer andtherapeutic agent. The coating may have s Mw, for example between 50,000to 300,000 g/mole. An exemplary embodiment is a UHMW PLLA scaffoldingwith a PDLA coating.

In further embodiments, a UHMW polymer scaffolding can have atherapeutic agent dispersed within and throughout the scaffoldingpolymer. The agent may be released into a vessel through diffusion fromthe scaffolding. Alternatively or additionally, the agent may bereleased due to erosion of the scaffolding. The agent, for example, canbe an anti-inflammatory substance that reduces or eliminatesinflammation to the vessel caused by degradation by-products of thescaffolding polymer. The scaffolding can alternatively or additionallyinclude an antiproliferative agent. Exemplary antiproliferative andanti-inflammatory agents include everolimus and clobetasol,respectively.

Solvent processing generally refers to forming a polymer construct suchas a tube from solution of polymer dissolved in a solvent. Inembodiments of the present invention, solvent processing methods can beapplied and adapted to making a bioabsorbable polymer construct in whichthe polymer is an UHMW polymer. Examples of solvent processing methodsinclude spray coating, gel extrusion, supercritical fluid extrusion,roll coating and dip coating.

Due to the high viscosity of the UHMW polymer, traditional meltextrusion is not possible because the temperatures and pressuresrequired cause excessive degradation of the polymer. However, some meltprocessing is may still possible without excessive degradation includingram extrusion and compression molding. However, the solvent processingtechniques described below are preferred for making stents.

The embodiments include applying gel extrusion to fabricate a constructmade from a UHMW polymer. In these embodiments, a polymer solutioncontaining a UHMW polymer dissolved in a solvent is conveyed through anextruder. Gel extrusion is described in patent application Ser. No.11/345,073 and is incorporated by reference herein in its entirety. Theembodiments described therein can be applied to the UHMW polymers. Ingel extrusion, the polymer solution has a viscosity low enough to beextruded at temperatures below the melting point of the polymer. Theserelatively low temperatures cause no or substantially no molecularweight degradation during the extrusion process. The solution isconveyed through a die at an end of the extruder to form a polymerconstruct such as tube. The solvent is removed from the polymer as itexits the die by quenching the exiting polymer, for example, in a waterbath.

Spray coating typically involves mounting or disposing construct such asa tube onto a support, followed by spraying a coating material (such asa polymer solution) from a nozzle onto the mounted stent. A sprayapparatus, such as EFD 780S spray device with VALVEMATE 7040 controlsystem (manufactured by EFD Inc., East Providence, R.I., can be used toapply a coating material to the construct. An EFD 780S spray device isan air-assisted external mixing atomizer. The coating material isatomized into small droplets by air and uniformly applied to theconstruct surface. Other types of spray applicators, includingair-assisted internal mixing atomizers and ultrasonic applicators, canalso be used for the application of the coating material. To facilitateuniform and complete coverage of the construct during the application ofthe composition, a tube can be rotated about the tube's centrallongitudinal axis. The tube can also be moved in a linear directionalong the same axis.

After application of the solution by spraying, the solvent is removedfrom the applied solution which leaves a layer of polymer on the surfaceof the construct. Drying or solvent removal depends on the volatility ofthe particular solvent employed, the solvent can evaporate essentiallyupon contact with the stent. Alternatively, the solvent can be removedby subjecting the coated stent to various drying processes, describedbelow.

In embodiments of the present invention, a tubular coating layer can beformed over a mandrel by spray coating with a solution of an UHMWpolymer. The tubular coating layer can then be removed from the mandreland a stent can be formed from the tubular coating layer. A tubularcoating layer of a desired thickness can be obtained by multiplerepetitions of spraying and drying.

Dip coating is a method of forming a coating layer onto an object whichincludes immersing the object in a coating material or solution thatincludes a polymer dissolved in a solvent, withdrawing the object fromthe solution, and removing solvent from the solution retained on thesurface of the object. Upon removal of the solvent, a layer of polymeris formed on the surface of the object. The steps above can be repeatedto form multiple layers of polymer over the object to obtain a desiredthickness of a coating layer.

In embodiments of the present invention, a polymer layer is formed overan object by immersing the object in a solution including an UHMWpolymer in a solvent. The object is withdrawn and solvent is removedfrom the UHMW polymer solution retained on the object to form a UHMWcoating layer on the object. The steps can be repeated to form a coatinglayer over the object of a desired thickness.

The object can be a cylindrical member or mandrel over which a tubularcoating layer is formed. For both spray coating and dip coating, themandrel can be made of any material that is not soluble in the solventof the polymer solution. In some embodiments, the mandrel is made of ametal such as aluminum or stainless steel. In other embodiments, themandrel is made from a glass with a polished surface. In some otherembodiments, the mandrel is made of a soluble material that is insolublein the solvent used for the coating. The coating layer is formed so thatits radial thickness or the thickness of the wall of the tubular layeris the desired thickness of a stent scaffolding. The coating layer maythen be removed from the mandrel and machined to form a stentscaffolding.

FIGS. 5A-C illustrate the dip coating process of the present invention.As shown in FIG. 5A, a mandrel 202 is lowered, as shown by an arrow 206into a container 204 having a solution 200 that includes a UHMW polymerdissolved in a solvent. The cylindrical axis of the mandrel isperpendicular to the surface of the solution, although the mandrel canbe immersed at an angle different from 90° to the solution surface. Asshown in FIG. 5B, the mandrel remains immersed in solvent 200 for aselected time or dwell time. Referring to FIG. 5C, mandrel 202 is thenremoved from solvent 200 as shown by an arrow 212. The cylindrical axisof mandrel 202 is perpendicular to the surface of the solution, althoughthe mandrel can be removed at angle different from 90° to the solutionsurface. The use of a 90° angle is expected to facilitate uniformity inthe coating thickness. Solution 210 is retained on mandrel 202 afterremoval from the solution 200 in container 204. Solvent is then removedfrom the retained solution 210 which results in the formation of acoating layer of the UHMW polymer. The solvent can be removed usingvarious types of drying methods described below.

Other dip coating processes can be envisioned by those skilled in theart. These include immersing only a small part of the mandrel into thesolution and while rotating parallel to the solution. This process helpsensure an even coating thickness.

In another embodiment, a hollow mandrel is dipped into the solution anda vacuum is drawn at one end of the mandrel causing the solution to bedrawn into the mandrel. When the mandrel is lifted from the solution,the solution will drain from the inside leaving the inside to themandrel coated with the polymer.

If the coating layer is a desired thickness, the coating layer can beremoved after solvent removal and machined to form stent. Alternatively,the steps in FIGS. 5A-C can be repeated one or more times until adesired thickness of polymer is achieved. In some embodiments, thecoated tube can be rotated 180° before each coating step is repeatedbecause gravity causes a greater volume of retained solution near alower end of the mandrel after removal of the mandrel from the solution.FIGS. 6A-B depict radial and axial cross-sections, respectively, of acoated mandrel 220 that shows mandrel 202 with a polymer coating layer216 with a thickness Tc.

There are several parameters in the dip coating process that can affectthe quality and uniformity of the coating layer. It is desirable for thetubular coating layer to be uniform circumferentially and along thecylindrical axis. Parameters include the concentration and viscosity ofthe polymer solution, the dwell time in solution, and the rate ofremoval of the mandrel from solution.

In some embodiments, polymer concentration can be at or near asaturation concentration. Such concentration is expected to result inthe highest viscosity and the thickest coating layer per immersion.Alternatively, polymer concentration can be less than saturation, forexample, less than 50% or less than 25% saturation. A more dilute andless viscous solution may result in a coating layer. However, a moredilute solution will require a higher number of repeated coating stepsto provide a final desired coating thickness.

The rate of removal of the mandrel from the solution can influence theuniformity of a coating layer of a single coating step and theconsistency of thickness of coating layers deposited in separate steps.The inventor has found in the removal time ranges considered, the rateof removal is directly proportional to the uniformity of coating layerthickness along the cylindrical axis for a coating from a single step.As the rate decreases, there was a greater difference in coatingthickness between the top end and bottom end of the mandrel. Theinventor has also found that the removal rate is directly proportionalto the consistency in thickness between coating layers deposited inseparate steps.

A stent body or scaffolding with layers of different material, as shown(FIGS. 3A-B) and described above, can be formed from a tube having awall with concentric layers of different material, with at least onelayer being a UHMW polymer layer. Such a layer can be formed in severalways. In one method, both the UHMW polymer and the low molecular weightpolymer layers can be formed by dip coating, as described above. Forexample, a UHMW layer of a desired thickness can be formed over amandrel and then a low molecular weight polymer layer can be formed of adesired thickness over the UHMW polymer layer by dip coating or by spraycoating. In another approach, a UHMW polymer layer may be formed by dipcoating over a low molecular weight polymer tube formed from extrusionor injection molding. The low molecular weight polymer tube may or maynot be disposed over a mandrel. When a mandrel is not used, thisapproach eliminates the step of removal of the dip coated tubularcoating layer from a mandrel.

The solvent can be removed from the solution retained on the mandrel bymethods known in the art including air drying or baking in an oven. Inair drying a gas stream is directed on or blown onto the mandrel. Thegas can be at room temperature or heated to increase the removal rate.

There are various ways to remove the tubular coating layer from themandrel to further process the coating layer in the fabrication of astent.

The coating layer can be formed over a mandrel made of a dissolvablematerial to be used as a dipping mandrel. After forming the coatinglayer, the mandrel can be dissolved by a solvent for the mandrelmaterial, but that is a non-solvent for the coating polymer. In anexemplary embodiment, the mandrel is a wax and the coating polymer isPLLA.

In another method of removal, the tubular coating layer is formed over ahollow mandrel or pipe with one end of the pipe covered by the coatinglayer. The polymer is formed such that it wraps around the ends of thepipe, creating a seal. After completing the coating layer, the coatinglayer can be cut off one end of the polymer wrapped pipe. Compressed airblow into the open end forces the tube off the mandrel.

In another method, the coating layer is laser machined to form the stentpattern while still mounted on the mandrel.

Another method includes forming the coating layer over an inflatedtubular balloon. The inflated tubular balloon is dip-coated as describedabove. Create a small scale balloon that is 3.2 mm inflated OD anddip-coat the balloon directly. After dip-coating, the balloon isdeflated and removed from the coating layer.

In another removal method, the mandrel is heated after forming thecoating layer. The heating is expected to loosen the coating layer,allowing it to be slipped off.

Another removal method includes application of oily or greasy coatingover the mandrel before dip coating. Once dip coating is completed, thecoating layer is slipped off.

In another method, a flexible rubber sleeve is wrapped around themandrel prior to dip coating. After coating is complete, the tube may bepulled off the mandrel by the sleeve. The tube is then removed from thesleeve.

In another method, after dip coating over a metal mandrel, the metalmandrel is cooled sufficiently to cause shrinkage of the mandrel,allowing the coating layer to be pulled off.

In another method, a mechanical slider may be used to force the tube offof the mandrel.

An automated dip coating system can include a syringe pump that performsa controlled immersion into a polymer solution, dwell time in thepolymer solution, and removal from the polymer solution of one or moremandrels. A syringe pump is a device designed to advance the plunger ofa syringe at a consistent, precise rate for continuously controlledliquid delivery. A specially adapted mounting system for mandrels can becoupled to the plunger. The motion of the plunger is designed to providecontrolled motion that immerses the mandrels at a controlled rate, toallow the mandrels to dwell in the solution for a selected time, and toprovide controlled motion that removes the mandrels from the solution ata controlled rate. An exemplary syringe pump for automated dip coatingis a Harvard Apparatus PHD 2000 programmable syringe pump.

FIG. 7 depicts a mandrel mounting disk 300 having a plurality of holesconfigured to hold mandrels for a dip coating operation. A plurality ofmandrels 304 are mounted on mounting disk 300 within the holes. FIG. 8Adepicts a system 310 for controlled dip coating of mandrels 304 mountedon mounting disk 300. System 310 has a syringe pump 320 positionedvertically and supported by a bracket 322. Syringe pump 320 includes asyringe plunger 324 that is coupled to mounting disk 300 on which aremounted a plurality of mandrels 304 (as illustrated in FIG. 7). In FIG.8A, mounting disk 300 is positioned such that mandrels are immersed in asolvent within a container 328. Plunger 324 is configured to movedownward, as shown by an arrow 330, at a controlled rate to immerse themandrels in the solvent and then allow the mandrels to dwell in thesolvent for a selected amount of time. FIG. 8B shows mounting disk 300and mounted mandrels 304 removed from the solution. Plunger 324 isconfigured to move upward, as shown by an arrow 332, at a controlledrate to remove the mandrels from the solvent and is further configuredto allow the mandrels to remain removed for a period of time to allowfor removal of solvent from the solution retained on the mandrels.

One advantage of using the UHMW for a stent body is that a stent bodymachined from the tube as-formed has sufficient mechanical properties tosupport a bodily lumen without a radial expansion step to improveproperties such as radial strength, modulus, and fracture toughness. Insome embodiments, the radial expansion step could be eliminated asadequate mechanical properties can be achieved without the improvementthat can be achieved as a result of radial expansion. in furtherembodiments, a UHMW polymer tube formed by methods described above canbe radial expanded prior to laser machining to form a stent. In theseembodiments, the stent can be formed from the expanded tube. The radialexpansion tends to increase the radial strength and toughness of thepolymer. The increase in strength is believed to be due thecircumferential polymer chain orientation and an increase incrystallinity, both induced by the expansion. The radial expansion canbe accomplished by a blow molding process. In such a process, thepolymer tube is disposed within a cylindrical mold with a diametergreater than the polymer tube. The polymer tube is heated, preferably sothat its temperature is above its Tg. The pressure inside of the tube isincreased to cause radial expansion of the tube so the outside surfaceof the tube conforms to the inside surface of the mold. The polymer tubeis than cooled below Tg and further processing steps can then beperformed, such as laser machining.

The degree of radial expansion or deformation may be quantified bypercent radial expansion:

$\left\lbrack {\frac{{Outside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {outside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Tube}} - 1} \right\rbrack \times 100\%$

In some embodiments, percent radial expansion can be 200-500%. In anexemplary embodiment, the percent radial expansion is about 300%.Similarly, the degree of axial deformation may be quantified by thepercent axial elongation:

$\left\lbrack {\frac{{Length}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}} - 1} \right\rbrack \times 100\%$

In some embodiments, the tube can be elongated before during or afterthe radial expansion. The percent axial elongation can be 30-200%.

For the purposes of the present invention, the following terms anddefinitions apply:

“Compression molding” is a method of molding in which the moldingmaterial, generally preheated, is first placed in an open, heated moldcavity. The mold is closed with a top force or plug member, pressure isapplied to force the material into contact with all mold areas, and heatand pressure are maintained until the molding material has cured. Theprocess employs thermosetting resins in a partially cured stage, eitherin the form of granules, putty-like masses, or preforms.

“Ram extrusion” refers to a process in which a resin is fed from ahopper and packed into a cylinder in repeated increments by areciprocating plunger. The frequency and amplitude of the plunger strokecan be controlled by an oil hydraulic system. The compressed materialmoves through a heated zone where it is fused into a profile matchingthe cross section of the barrel or die. The output rate is proportionalto the length and frequency of the ram strokes. Die length, electricalheater capacity, hydraulic system power and maximum force, and thestrength of the construction materials determine equipment capability.

“Molecular weight” can refer to the molecular weight of individualsegments, blocks, or polymer chains. “Molecular weight” can also referto weight average molecular weight or number average molecular weight oftypes of segments, blocks, or polymer chains.

The number average molecular weight (Mn) is the common, mean, average ofthe molecular weights of the individual segments, blocks, or polymerchains. It is determined by measuring the molecular weight of N polymermolecules, summing the weights, and dividing by N:

${\overset{\_}{M}}_{n} = \frac{\sum_{i}{N_{i}M_{i}}}{\sum_{i}N_{i}}$

where Ni is the number of polymer molecules with molecular weight Mi.The weight average molecular weight is given by

${\overset{\_}{M}}_{w} = \frac{\sum_{i}{N_{i}M_{i}^{2}}}{\sum_{i}{N_{i}M_{i}}}$

where Ni is the number of molecules of molecular weight Mi.

The “inherent viscosity” (of a polymer) is the ratio of the naturallogarithm of the relative viscosity, ηr, to the mass concentration ofthe polymer, c, i.e. ηinh=(ln ηr)/c, where the relative viscosity (ηr)is the ratio of the viscosity of a polymer solution, η, to the viscosityof the solvent (ηs), ηr=η/ηs.

“Ambient temperature” can be any temperature including and between 20°C. and 30° C.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness in this caseare in energy per unit volume of material. See, e.g., L. H. Van Vlack,“Elements of Materials Science and Engineering,” pp. 270-271,Addison-Wesley (Reading, Pa., 1989).

The underlying structure or substrate of an implantable medical device,such as a stent can be completely or at least in part made from abiodegradable polymer or combination of biodegradable polymers, abiostable polymer or combination of biostable polymers, or a combinationof biodegradable and biostable polymers. Additionally, a polymer-basedcoating for a surface of a device can be a biodegradable polymer orcombination of biodegradable polymers, a biostable polymer orcombination of biostable polymers, or a combination of biodegradable andbiostable polymers.

It is understood that after the process of degradation, erosion,absorption, and/or resorption has been completed, no part of the stentwill remain or in the case of coating applications on a biostablescaffolding, no polymer will remain on the device. In some embodiments,very negligible traces or residue may be left behind. For stents madefrom a biodegradable polymer, the stent is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.

Representative examples of polymers that may be used to fabricate animplantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(trimethylene carbonate), polyester amide,poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters)(e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin,fibrinogen, cellulose, starch, collagen and hyaluronic acid),polyurethanes, silicones, polyesters, polyolefins, polyisobutylene andethylene-alphaolefin copolymers, acrylic polymers and copolymers otherthan polyacrylates, vinyl halide polymers and copolymers (such aspolyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether),polyvinylidene halides (such as polyvinylidene chloride),polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such aspolystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Another type of polymer based on poly(lacticacid) that can be used includes graft copolymers, and block copolymers,such as AB block-copolymers (“diblock-copolymers”) or ABAblock-copolymers (“triblock-copolymers”), or mixtures thereof.

Additional representative examples of polymers that may be especiallywell suited for use in fabricating or coating an implantable medicaldevice include ethylene vinyl alcohol copolymer (commonly known by thegeneric name EVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

EXAMPLES

Some embodiments of the present invention are illustrated by thefollowing examples. The examples are being given by way of illustrationonly and not by way of limitation. The parameters and data are not beconstrued to unduly limit the scope of the embodiments of the invention.

Example 1 Exemplary UHMW Resin for Use in Making Stent

A UHMW polylactide resin with brand name PURASORB PL and batch no.0404002128 was obtained from PURAC located in the Netherlands. Theproperties of the resin are:

Appearance: white to light tan granules Inherent Viscosity: 8.36 dl/g(chloroform, 25° C., c = 0.1 g/dl) Specific rotation (chloroform 25° C.)−157.8 Melting range (DSC, 10° C./mon) 177.6-192.3° C. Heat of fusion(DSC, 10° C./mon) 74.1 J/g Residual Solvent <0.01% Residual monomer <0.1%

Example 2

Illustration of Dip Coating Method with PLLA, Mw=2,200,000 g/Mole

PLLA tubes were made by dip coating with a solution of PLLA dissolved inchloroform. The concentration of the solution was a 2.5% weight byvolume sample which provided uniform, complete dissolution.

80 mL of solution was prepared by dissolving PLLA using 78 mL ofchloroform (Acros Organics: Chloroform, Extra Dry, Water <50 ppm,Fischer Science Catalogue number AC32682-0010) and 2.5 g of PLLAgranules received from PURAC biochem, Gorinchem, The Netherlands. Thesolution was dissolved overnight in a closed container within a fumehood.

Once the solution was prepared, a dipping apparatus was set up with ametal mandrel of 0.125 in diameter. Mandrels were lowered by hand intothe solution and removed manually at an approximate rate of 1, 2, or 3cm per second. The solvent was then allowed to evaporate.

Coating thickness was determined by caliper measurement after each coatper the following formula: Tw=(Ti−Tm)/2, where Tw is the wall thickness,Ti is the total thickness (measured) and Tm is the fixed mandrelthickness (0.125 in). This process was continued until appropriate wallthickness was obtained, 0.006 in. Tubes were made using three differentremoval times from the solution: 15 seconds, 30 seconds, and 45 seconds.

The results from the dip coating are shown in FIG. 9. The averagecoating thickness of each of the three coats is shown. A longer removaltime corresponds to a slower speed of mandrel removal. As shown in FIG.9, the runs of fastest removal time resulted in the most linearrelationship of thickness along the length of the tube. A slowervelocity of mandrel removal from the solution resulted in a greaterdifference in thickness at the bottom of the tube compared to othersections. This could be due to the greater amount of time allowed to thepolymer solution to flow down the length of the tube due to the pull ofgravity.

Example 3 Exemplary Dip Coating Process

-   1. Mix 195 ml chloroform and 6.25 g polymer to create 200 ml of    solution (2.5 wt %) need to mix solution until all granules have    dissolved (using stir bar with stir plate at 250 rpm)-   2. Fill up dip container with solution to the marked line-   3. Place hollow mandrels (0.125″) into fixture-   4. Dip mandrels for 30 seconds, let dry for 30 minutes-   5. Invert mandrel and repeat 3 more times (total 4 dips, invert    mandrels after every dip)-   6. Put tubes into vacuum at room temperature for 48 hours-   7. Cut the tubes to 110 mm and remove the tubes from the mandrels    using the mechanical slider and place the tubes on the solid    mandrels (0.123″)-   8. Put tubes back into vacuum for 24 hours-   9. Cut tubes to 100 mm in length-   10. After 24 hours, measure the proximal, middle, and distal ID and    OD-   11. Remove tubes from mandrels

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A method of making a stent body for supporting avascular lumen, comprising immersing a cylindrical member in a solutionincluding a bioabsorbable polymer dissolved in a fluid, wherein thebioabsorbable polymer has an inherent viscosity greater than 6 dl/g orhas a weight average molecular weight greater than 1 million g/mole,removing the member from the solution, wherein a portion of the solutionremains on the surface of the member upon removal from the solution;removing solvent from the solution remaining on the member to form atubular layer of the bioabsorbable polymer on the member; repeating theimmersion step, removal from the solution step, and removal of thesolvent step to form a final tubular layer of bioabsorbable polymer onthe member of a desired thickness; and forming a stent body from thefinal tubular layer.
 2. The method claim 1, wherein the member isremoved from the solution in less than 30 seconds.
 3. The method claim1, wherein the member is immersed with its cylindrical axisperpendicular to the surface of the solution.
 4. The method claim 1,wherein the member is rotated 180° prior to repetition of the immersionstep.
 5. The method claim 1, further comprising radially expanding thefinal tubular layer and forming the stent body from the expanded tube.6. The method claim 1, wherein the solution further comprises atherapeutic agent, and the final tubular layer comprises some of thetherapeutic agent.